Device and system for improved imaging in nuclear medicine and mammography

ABSTRACT

A method and apparatus for detecting radiation including x-ray, gamma ray, and particle radiation for radiographic imaging, and nuclear medicine and x-ray mammography in particular, and material composition analysis are described. A detection system employs fixed or configurable arrays of one or more detector modules comprising detector arrays which may be electronically manipulated through a computer system. The detection system, by providing the ability for electronic manipulation, permits adaptive imaging. Detector array configurations include familiar geometries, including slit, slot, plane, open box, and ring configurations, and customized configurations, including wearable detector arrays, that are customized to the shape of the patient. Conventional, such as attenuating, rigid geometry, and unconventional collimators, such as x-ray optic, configurable, Compton scatter modules, can be selectively employed with detector modules and radiation sources. The components of the imaging chain can be calibrated or corrected using processes of the invention. X-ray mammography and scintimammography are enhanced by utilizing sectional compression and related imaging techniques.

FIELD OF THE INVENTION

This invention relates to an improved system for radiographic imagingand material analysis and more specifically for nuclear medicine andmammography imaging.

BACKGROUND OF THE INVENTION

Two general imaging problems in radiology involve the determination of aradiation source distribution and/or the effect of a filter, in thiscase a patient, on the radiation source distribution. Consider the fieldof nuclear medicine where the radiation source or other radionuclidedistribution emits photons or positrons, Image data acquisition innuclear medicine presents several challenges in addition to constraintsimposed by finite acquisition times and patient exposure restrictions.Most photon energies that are of interest in nuclear medicine are higherthan the typical photon energies employed in diagnostic x-rayradiography. In particular, Positron Emission Tomography (PET) involvesthe detection of pairs of very high energy photons due to annihilationevents Unfortunately, the photon radiation source, such as aradionuclide, used in nuclear medicine is not directional and the sourcedistribution within the body is not precisely known.

Photons that escape the body may be scattered, altering their energiesand/or direction vectors. It is desirable for many applications todiscriminate against scatter radiation reaching the detector based onenergy and/or direction. It may also be desirable to only detectradiation with a specific direction vector, since many detection systemspossess poor directional discrimination capability and have finiteresponse times within which to detect events, thereby limiting detectionrates. Thus detection systems used in nuclear medicine such as Gammacameras or PET scanners often employ conventional, such as attenuatingor rigid geometry, focused or unfocused collimators, often referred toas grids or grid collimators, to help define the direction vectors of adetected photons. The direction vectors and energies of non-scatteredphotons are well-defined. Unfortunately, the emission of photons fromthe source distribution is non-directional and the radiation sourcedistribution itself is typically not well-defined. A Compton-scatteredphoton suffers an energy loss and change in direction vector whereas acoherent or Rayleigh scattered photon only has its direction vectoraltered. In general x-ray radiography the source is a x-ray tube,although a radionuclide maybe substituted, used in a point, slit, slot,or area imaging configuration. The energy distribution and directionvector of the radiation from a x-ray tube are approximately known. Theseparameters are typically well-defined for a collimated radionuclidesource used in an application such as point-scan Compton scatter imagingand material analysis. A number of detection formats are in usedepending on the application. A planar detector geometry is typicallyutilized for applications such as mammography, angiography, and chestradiography which typically employ detectors such as x-ray film-screendevices, or storage phosphor screens, or image intensifiers coupled tocameras. Slit- and slot-scan formats are also available, usuallyincorporating improvements to the detectors and, in some instances, theradiation source. Additional image acquisition formats includering-shaped detectors or flat detectors for fan-beam or cone-beamtomography, respectively. Common detector geometries used in nuclearmedicine typically include one or more planar detectors, which arebasically standard Gamma cameras, with attached conventional collimatorsor ring detectors, used in Positron Emission Tomography. Imaging systemsbased on standard Gamma camera and related detector designs arefrequently used for a number of nuclear medicine studies such as heart,brain, thyroid, gastro-intestinal, whole body, and breast imaging,including scintimammography. A basic Gamma camera design employs alarge, planar array of scintillation crystals or a single, large, planarscintillation crystal optically coupled to an array of photomultipliertubes (PMTs). A conventional focused or unfocused collimator istypically mounted to the face of the Gamma camera. This inflexibleimaging system is then positioned such that the region of interestcontaining the source distribution is within the field of view. Itprovides a limited degree of spatial resolution and energy resolutionwhile removing some fraction of unscattered and scattered radiation thatwould otherwise degrade image quality. Unfortunately a substantialfraction of useful unscattered radiation is also attenuated. Anotherinfrequently used design replaces the conventional collimator with acoded aperture such as a uniformly redundant array aperture which isalso based on photon attenuation and is typically rigid. Commercialsystems may use one, two, or three Gamma camera detector units. Onecommercial system eliminates the use of scintillator crystals and PMTswith a rigid, planar, 2-D CdZnTe semiconductor detector manufactured byabutting four 2-D CdZnTe arrays of moderate size. Techniques forabutting 2-D silicon arrays are well-known in the art. Drawbacks toemploying large- or medium-sized 2-D CdZeTe arrays capable of highdetection efficiency include the difficulty of growing thick CdZnTecrystals with acceptable levels of defects and creating a low noise, 2-Darray readout structure on top of a large- or medium-size CdZnTecrystal. Grid collimators are still desirable for many applicationssince the direction vectors of detected photons are otherwise poorlydefined. A design which replaces a conventional collimator with arelatively thin, planar semiconductor, often Ge, array of moderate size,which serves as a Compton scatterer is referred to as a Comptonelectronic Gamma camera. This system is still being refined. Thedetector module array described below can be used in place of a standardGamma camera in a Compton Gamma camera system.

Nuclear medicine imaging applications are complicated by the fact thatthe spatial distribution of the source within a region of the patient ispoorly defined. One way to simplify this problem is to use emittedphotons of known energies. For example, a source that has one or moreemission energies of a narrow energy bandwidth may be utilized. Theproblem now is the reconstruction of the source distribution rather thanthe calibration of the source distribution. The measured sourcedistribution, i.e., the apparent source distribution, represents thefiltered true source distribution, assuming self-attenuation is small.In certain nuclear medicine applications estimates of the true sourcedistribution are obtained by calibrating the contribution of the filter,which may be the patient, to the apparent source distribution. Photontransmission measurements are made in order to estimate the effect oftissue scattering and absorption or attenuation on radiation sourcemeasurements by using a reference source that is external to thepatient. Unfortunately, measuring photon transmission through the bodydoes not duplicate the actual imaging chain acquisition format used innuclear medicine where photon are transmitted out of the body. Photonsin the two instances do not traverse comparable paths.

SUMMARY OF THE INVENTION

In accordance with the present invention, a radiation detectionapparatus is provided for radiographic imaging and material compositionanalysis in which the apparatus can dynamically configure its arraygeometry and radiation detector parameters for a specific imaging taskor it can use an existing radiation detection geometry and settings.This invention is particularly suited for x-ray and gamma ray imaging innuclear medicine, including scintimammography, and x-ray radiography,specifically, x-ray mammography. There are several advantages inherentto this invention. Superior detectors in cost-effective formats can beutilized and detectors with different properties, including materials,resolution, response time and noise characteristics, can be used withinan array. One or more radiation detectors are incorporated into adetector module and one or more modules make up a detector module array.The detector modules transmit detected photon image data and relevantmodule parameters to a computer system which utilizes this informationto electronically-control the modules and in some cases attachedcollimators. This system is implemented using detector sub-arrays,comprised of one or more detector modules, and detector arrays in orderto enhance image quality or analysis capability. Conventionalattenuating or rigid geometry collimators, including ones characterizedby coded apertures, and unconventional, including x-ray optic,configurable (adaptive), and Compton scatter module, collimators can beemployed to improve the energy and/or spatial resolution for the photonradiation detection system. In a similar manner additional types ofradiation optic collimators such as neutron optic collimators orelectron optic collimators capable of focusing electric or magneticfields, can be used with neutrons or charged particles, respectively.

In a preferred embodiment semiconductor detectors with appropriategeometries, such as edge-on detectors; thick, linear array detectors; orsmall, thick, 2-D array detectors, are incorporated into detectormodules which are mounted within a frame and configured as an array ofdetector modules. Detector modules contain one or more detectors,possibly with different properties. A detector array contains one ormore modules or types of modules. For nuclear medicine imagingapplications detector sub-arrays, comprised of one or more modules, orthe entire detector array can be positioned and oriented with respect tothe radiation source by an operator or by direct computer control.Collimators and shielding can be attached to or integrated into themodule, including interfacing with module electronics if appropriate.Modules communicate with the computer system which monitor and controlmodule and collimator parameters and collect and process radiation datarecorded by the detectors. Modules may communicate directly or through ashared network with the computer system. Computer-controlled servicesinclude sending electronic instructions to the module mounting hardware,the module, and the collimators, if appropriate. Electronic instructionscan initiate actions such as detector array motion, adjustment of therelative position or orientation of one module with respect to othermodules, manipulation of a collimator, and the modification of moduleoperating parameters, such as detector signal amplification, filtering,resolution, temperature, operating voltage or sampling rate. Sincepositioning machinery can be incorporated into the module, actuators canbe employed to adjust the position and orientation of the detector. Theactuators can also manipulate the positions and orientations ofappropriate collimators. A novel collimator design utilizes actuators toalter the configuration of a collimator. The computer-based monitor andcontrol capabilities can be used to track and adjust the locations ofmodules while they are in motion. Positions, orientations, and motion ofall detectors and relevant collimators are recorded and updated asneeded throughout the image acquisition process.

A typical nuclear medicine imaging session begins with an operatorselecting from a computer display menu a specific detection system withpre-defined array geometry, collimator, and module settings appropriatefor the desired imaging task. The detector array configuration canalready exist or it can be set up by the computer system. Once abaseline detection system is established, an operator can then adjustand fine tune the detector array position and settings or leave thedetector array adjustments and tuning under computer control. Whileunder computer control electronic instructions can be issued dynamicallyin response to detector module parameter values and detected radiationdata that is transferred to the computer system for processing, display,and storage during image acquisition or adaptive imaging. Electroniccommands can be used to control the array geometry and motion, detectormodule parameters, and some types of collimators. Thus an informationfeedback loop can be implemented as a means of tuning detection systemparameters. For some imaging or analysis applications it will besufficient to configure the detector array based on either a standardgeometry, such as line, plane, open box, wedge, ring, cylinder, ellipse,ellipsoid, or sphere, or a contoured geometry, in order to compensatefor the radionuclide distribution within the subject and/or the shape ofthe subject at the region of interest. For example, configurations maybe based on the breast size of a woman or on the head size, waist size,or chest size of an adult, child, or infant. A versatile design allowsat least a subset of these detector array geometries to be generated “onthe fly”. A less-versatile design still utilizes modules, but themodules are fixed within a specific detector array geometry or they areconstrained to move to specific positions, for example, along a track,within a specific detector array geometry. Less-versatile designs reducethe mechanical complexity of the detection system and may be sufficientfor specific imaging tasks. An optional capability is to allow theentire array to undergo discrete or uniform motion. The simplest exampleof this capability would be to scan a radiation source with a detectorarray comprised of a single detector module.

In another embodiment, semiconductor detectors are replaced by othertypes of suitable detectors, such as scintillation detectors, gasdetectors, liquid detectors, or superconducting detectors.

In another embodiment reference sources are introduced into the subjectand then imaged. The size, shape, intensity, and emission spectrum ofthe reference sources are known. This allows measurements to be made ofphoton attenuation due to material in the photon path prior to reachingthe detector. This information can be used to estimate the true sourcedistribution from measurements of the apparent source distribution madeduring image acquisition in a nuclear medicine test. The referencesource can also be used to focus the detector array in order to tune theimaging chain.

In another embodiment detector modules and collimators are incorporatedinto x-ray radiography slit scan imaging systems. X-ray opticcollimators can be used to increase the intensity and modify thespectrum of the x-ray radiation that is recorded by the detector module.A single x-ray source is combined with a x-ray optic collimator and ax-ray detector module and used for a x-ray mammography slit scan system.Another improvement involves aggressively compressing sections of thebreast and acquiring separate images of the highly-compressed sectionsrather than acquiring a single image of the entire, mildly-compressedbreast.

The system of the present invention may utilize devices detailed inprior inventions for slit-scan or slot-scan radiographic x-ray imagingin which photons are detected directly using edge-on array detectors;small, 2-D semiconductor array detectors; or semiconductor arraydetectors coupled to scintillators. This new device can also use thick,linear semiconductor array detectors and thick, small, 2-D semiconductorarray detectors in addition to other types of detectors. Manufacturingcosts for these detectors are much less than those associated withlarge-area or moderate-area, thick, planar, 2-D semiconductor arraydetectors made from materials such as, but not limited to, CdZnTe, CdTe,GaAs, Ge, Si, SiC, or HgI2. The detector format is also compatible withdetectors such as thin, linear semiconductor arrays or thin, small, 2-Dsemiconductor arrays coupled to scintillators. For example, thin, linearsemiconductor arrays of avalanche photodiodes coupled to scintillatorscan be used as radiation detectors. This approach can be extended toinclude scintillators coupled to integrated photoemissive cathodes orsmall PMTs; small, gas microcapillary detector assemblies; or smallsuperconducting array detectors. Consider a scenario in which radiationis incident upon a planar edge-on detector. The detector thickness(height) now defines the maximum detector entrance aperture while thelength or width of the detector area now defines the maximum attenuationdistance for edge-on radiation detector designs including semiconductordrift chamber, single-sided strip, and double-sided strip detectors,including micro-strip detector versions. Strip widths can be tapered orcurved, in the case of drift chamber detectors, if focusing is desired.In the case of double-sided parallel strip detectors in which opposingstrips are parallel, both electrons and holes can be collected toprovide 2-D position information across the aperture. If strips on oneside run perpendicular to those on the other side, thendepth-of-interaction information can be obtained. If strips aresegmented in either a single-sided or double-sided parallel stripdetector then depth-of-interaction information can be obtained andreadout rates can be improved.

These and other advantages of the present invention will become apparentupon reference to the accompanying drawings and the followingdescription.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 illustrates a perspective view of a detector module array.

FIGS. 2 a–2 c and 2 d(i)–(iii) illustrate perspective views of variouscylindrical, spherical and wearable detector array geometries.

FIG. 3 a illustrates perspective views of box-shaped implementations ofdetector arrays.

FIG. 3 b illustrates an L-shaped detector array configuration.

FIG. 3 c illustrates a compliant detector array configuration.

FIG. 3 d–3 f illustrate configurable arrays operated by actuator arms.

FIG. 4 a–4 b illustrate perspective views of an edge-on detector modulewith an unconventional minifying/tapered capillary x-ray opticcollimator.

FIG. 5 a(i)–(ii) illustrate a configurable dual x-ray refractive lens.

FIG. 5 b(i)–(ii) illustrate nested refractive lenses.

FIG. 6 illustrates a perspective view of an electronically controlledconfigurable collimator.

FIG. 7 a illustrates an electronically controlled configurablerefractive lens with an associated configurable x-ray mirror andradiation source.

FIG. 7 b illustrates a fixed-focal length capillary x-ray lens with anassociated configurable x-ray mirror and radiation source.

FIG. 7 c–7 d illustrate an array of capillary x-ray optic lenses alignedwith an array of radiation sources in the form of a cylindrical anode.

FIG. 7 e illustrates a wedge-shaped capillary x-ray optic lens alignedwith a dense array of radiation sources.

FIG. 8 illustrates a planar detector array comprised of strip detectormodules positioned within a frame.

FIG. 9 a illustrates a capillary x-ray lens system designed with asingle gantry arm.

FIG. 9 b illustrates a multiple gantry arm system.

FIG. 9 c illustrates a system utilizing additional x-ray opticsintroduced between a capillary x-ray lens and compression plates.

FIG. 10 a–10 b illustrate a system incorporating flat and contouredcompression plates.

FIG. 10 c illustrates the use of overlapping images to ensure completeimaging of the breast.

DETAILED DESCRIPTION

General Detector Array

In one embodiment of the present invention, as illustrated in FIG. 1, adetector array 1000 preferably incorporates separate, discrete detectormodules 102, illustrated here as edge-on or strip/micro-strip detectors,configured in a planar geometry to optimize the detection of incidentradiation 107. The detector array 1000 may be utilized as part of agamma camera. Currently, gamma cameras are not based on detector arrayssuch as detector array 1000 which incorporates discrete modules 102.

Detector modules 102 utilize one or more detectors 101, typically arraydetectors, which can have different properties. Note that several of themodules 102 include more than one detector. Additionally, linear arrayor small, 2-D array semiconductor detectors may be incorporated into thedetector modules 102. Each module 102 also includes a base 106 and acommunications link 103.

The base 106 preferably contains detector electronics, power managementcomponents, temperature control components, and a data or informationchannel for communicating with the computer system. The base 106 mayalso incorporate a module electronic readout unit that includes a signalconditioner or filter, an amplifier, an analog-to-digital converter, anda communication interface. Additionally, the detector module 102 may becoupled to an electronically-controlled thermoelectric cooler or othertemperature regulating device which resides in the detector module base106. In this embodiment, the temperature-regulating device providestemperature control for the detector module 102 and its electronicreadout unit.

The communications link 103 provides power to the module 102 andconnects the base 106 to a computer system. Through the attachment withthe base 106, the link 103 enables a computer system to control theconfiguration of the module 102. The communication link 103 preferablyis used to off-load the digitized detector radiation data to a computersystem for analysis and image reconstruction. The computer system, whichcan include general purpose, dedicated, and embedded computers, providesmonitor and control services to modules 102 and to the entire detectorarray 1000. The computer system evaluates module, detector arrayparameters, and the detected radiation image data. The detected data isprocessed and can be displayed and stored if desired. Additionalrelevant module information, such as temperature, amplifier settings,detector voltages, position, orientation, and motion information, can betransmitted to this computer system over the communication link 103.Alternatively, a separate communication channel may be incorporated totransfer the additional information between the module 102 and thecomputer. The computer system transmits instructions that update thedetector array 1000. This establishes a dynamic information feedbackloop that is useful for adaptive imaging.

Each module 102 may optionally have its own radiation shielding 104 andcollimator 105 mounted on the wall of the module 102, although only onemodule is shown with these items for clarity. Module walls are typicallythin, which permits radiation-shielding 104 to be attached to the modulewall or inserted between adjacent modules 102 when needed. Asillustrated in FIG. 1, collimators 105 are placed in operable contactwith the detector modules 102. However, the array 1000 is capable ofoperating without collimators 105. Even in the absence of collimators105, collimation exists to a limited extent because the modules 102 arediscrete and physically separated. In this alternative embodiment, thearray 1000 is designed to detect incident radiation 107 using detectormodules 102 without attached collimators 105.

Turning back to the embodiment illustrated in FIG. 1, both conventionaland unconventional collimators 105 can be attached to the detectormodule 102. If the collimators 105 are capable of beingelectronically-controlled to perform mechanical alignment or tomanipulate unconventional collimators, then the module 102 may utilizethe communications link 103 to provide power management andcommunication capability to the collimator 105. The communication link103 is also used to transmit collimator parameters and settings betweenthe module 102 and the computer system.

A detector array 1000 can be comprised of more than one type of detectormodule 102. A number of array geometries, in addition to the standardplanar detector array format, can be utilized. This design may also beimplemented by using semiconductor detectors coupled to scintillators aswell as other types of detectors. These detectors are described inNelson, U.S. Pat. No. 4,560,882, which is incorporated by reference forall it discloses and describes. Limiting the focusing of the modules 102using edge-on strip detectors is possible by tapering the edge-onstrips, as described in Nelson, U.S. Pat. No. 4,937,453, filed May 6,1987, which is hereby incorporated by reference for all it discloses anddescribes.

The increased detector density is useful for enhancing the imaging ofselect regions of the subject. For example, the array 1000 illustratedin FIG. 1 is particularly suited to form large area, 2-D detectorarrays, such as those described by Nelson, U.S. Pat. No. 4,937,453, dueto the close proximity of multiple modules 102.

The detector array 1000 and individual detector modules 102 can bescanned or dithered, i.e. moved repeatedly between adjacent locations,as needed so as to provide appropriate sampling of spatial regions whichwould otherwise be ‘dead areas’ due to a lack of detectors at thosepositions. Scanning motion is suitable for sampling regions that wouldotherwise require multiple detector modules to fill. Dithering issuitable for sampling regions that are typically less than the size ofdetector pixels. A detector module 102 can incorporate more than onedetector 101 with the same or different properties. A detector array1000 can use more than one type of detector module 102. For example,detectors 101 or detector modules 102 with different energy resolutions,spatial resolutions, stopping powers, and readout rates can be combinedin order to match the imaging characteristics of the device to specificas well as general applications. Temperature management, typicallycooling, of the detectors 101 may be utilized as needed. Preferably thedetectors 101 will operate at or near room temperature for applicationsin nuclear medicine, but this may not always be possible. Several safetyfeatures can be included with the detector array 1000 to restrict itsspeed of motion and proximity to the subject. These can include opticaland acoustic range sensors as well as pressure-sensing devices.Computer-controlled positioning and sensor devices have been widely usedfor many years in applications such as medical imaging, robotics,factory automation, precision machine tools, micromachines, andaviation.

An external container (not shown) preferably is typically employed toshield the detector array 1000 and any electronic components fromexternal electromagnetic fields and physical contact.

Specialized Detector Arrays

The present invention includes specialized detector arrays such as thoseillustrated in FIG. 2 a–2 d. FIG. 2 a–2 d illustrate perspective viewsof fixed cylindrical-like (FIG. 2 a), spherical-like (FIG. 2 b), andwearable (FIG. 2 c) detector array geometries.

Cylindrical and Spherical Detector Arrays

In FIG. 2 a and 2 b, edge-on detector modules, 102(a) and 102(b), arecapable of motion along tracks, 110(a) and 110(b). Such a characteristicenables the modules 102(a) and 102(b) to sample gaps between individualmodules. Turning to FIG. 2 a, the modules 102(a) are shown with sideshielding 112 and a radiation source 111. The shielding 112 may beincluded on the modules 102(a) in order to minimize the detection ofradiation escaping from neighboring detectors. The modules 102(a) areconfigured to move in a cylindrical fashion along the tracks 110(a).

Turning to FIG. 2 b, the tracks 110(b) are configured to allow themodules 102(b) to move in a spherical motion with both up and down andside to side directionality possible. In both FIG. 2 a and FIG. 2 b, themodules, 102(a) and 102(b), move independently along tracks, 110(a) and110(b), as part of a detector array or detector sub-array, or both typesof motion can be executed in order to improve the sampling uniformity ofthe subject. The optional use of a flexible track (not shown) permitsthe detector modules, 102(a) and 102(b), to follow the contours of thesubject more closely by enabling the modules, 102(a) and 102(b), to movein conformity with the subject. Using a flexible track would allow for acontoured geometry.

The computer system keeps a chronological, real time record of moduleand array parameters, including position and orientation of detectorsand collimators, motion, detector amplification and noise, during imageacquisition and during detector array calibration.

Detector array configurations such as those shown in FIG. 2 a and FIG. 2b may be extended or filled-out to form a more complete cylinder, suchas a ring shape, or even spherical arrays of detectors. Detector arrayconfigurations may also be shaped to fit other standard geometries,including slit, slot, line, plane, open box, wedge, ellipse, ellipsoid,or a combination of geometries. Collimators can be employed withindividual edge-on detectors, such as illustrated in FIG. 1.

Wearable Detector Arrays

Turning now to FIG. 2 c, a wearable detector array 2000(a) in the formof a detector vest, similar in form to an upper torso “body armor”shell, a down-filled winter vest, or a breastplate, is illustrated. Thedetector vest 2000(a) is designed such that it approximately conforms tothe shape of the patient. Straps 123 may be included to allow thepatient to securely wear the detector vest 2000(a). Straps 123 may alsobe used with other wearable detector arrays. For example, a wearablehelmet could incorporate an adjustable chinstrap. The adjustable straps123 also permit the detector vest 2000(a) to be positioned at thedesired location for a range of body types.

A detector array 1000, sufficient in size for imaging the heart, isincorporated into the detector vest 2000(a). This configuration can becost-effective for particular cardiac imaging studies since the area ofthe detector array, and any attached collimators, is not much largerthan the projected area of the heart onto the plane of the detectorarray. An additional benefit is the reduction in weight that is possibleby employing a small detector array and collimator. A detector vest2000(a) such as that illustrated in FIG. 2 c enables effective samplingof the region of interest throughout the examination, mitigatingproblems due to undesired patient motion. Wearable detector devices mayalso permit claustrophobic patients to undergo testing since one ormultiple large gamma camera heads will not be positioned about thepatient so as to create a confining effect. Communication links 103 areused to facilitate communication between the detector array 1000 and acomputer and to supply power to the detector modules 102.

An optional support harness 122 reduces the additional stress imposedupon the patient that may be generated due to the weight and bulk of thedetector vest 2000(a). If the detector vest 2000(a) is too heavy orbulky for the patient then a flexible suspension system using thesupport harness 122 is employed to provide at least partial support forthe detector vest 2000(a). The use of a support harness 122 is verysimilar to the concept of using a training harness for a patient oranimal recovering from injuries or for simulating effects such asreduced gravity for an astronaut. A flexible suspension system can alsobe employed with other wearable detector devices if needed.

An alternate implementation of a wearable detector which still permitslimited mobility is a detector array configuration that is supported bya stand with adjustments for height, tilt, and rotation. The patientpresses the appropriate body part (head, neck, chest, stomach, etc.)against the detector array configuration while stationary or performingan exercise regimen. This imaging format maintains a reasonablealignment between the detector array and body region and is similar inprinciple to using exercise equipment such as cross-country ski trainerswhere an individual presses the hips or stomach against an adjustablepad.

Detector arrays can be incorporated into similar equipment to createwearable detector arrays such as wearable detector helmets/head gear,detector neck braces, detector brassieres, and detector belts orgirdles. FIG. 2 d(i)–2 d(iii) illustrate different wearable devices,namely a detector vest (2000(b), FIG. 2 d(i)), a detector helmet(2000(c), FIG. 2 d(ii)), and a detector brassiere (2000(d), FIG. 2d(iii)), that use open frames, 124(b)–124(d), in the shape of thedesired geometry for mounting detector modules and collimators. The openframes 124(b)–124(d) provide a grid for mounting the detector array 1000and any associated collimators to examine different sections of thesubject. For example, a detector array 1000 could be mounted on an openframe detector vest 2000(b) so as to image the heart and lungs andanother detector array 1000, possibly with modules 102 of differentproperties and collimators 105, could be mounted so as to image thekidneys and gall bladder. The many possible configurations of detectorarrays 1000 that can be implemented in a wearable device such as awearable detector vest enable it to provide the same types of views,including 180 degrees, 240 degrees, and 360 degrees, which single andmultiple head Gamma cameras are able to acquire.

Additionally, compressible pieces of foam or other expandable bladders125 may also be attached inside a wearable detector in order to allowfor a further customized fit to the patient. Bladders 125 are alsoutilized to minimize contact between the patient's body and rigid areasof the wearable detector, including the detector array 1000 withdetector modules 102 and collimators 105. The bladders 125 are directlycomparable to the use of foam and expandable bladders in athletic gearsuch as football helmets. If the pressure exerted by an expandablebladder 125 is modulated, an established technique that is used toassist circulation in the legs of bed-ridden patients, then specificphysiological studies that depend on circulation, including breastphysiology, can also be conducted.

If needed, the patient can be cooled by circulating air or a containedliquid between regions of the wearable detector and the patient's bodythat do not interfere with image acquisition. The power andcommunication connections, such as the communications links 103, thatinteract with the detector arrays 1000 can also interact with or controlcooling devices (not shown) and expandable bladders 125.

Detector arrays 1000 are capable of being installed in fixedconfigurations in wearable detectors or they can be designed to beremovable. For example, removable detector arrays 1000 may be configuredto snap into pre-defined positions, dynamically establishing power andcommunication connections with the power source and computer,respectively, and permitting customization. The locations of thedetector arrays 1000, and modules 102 within the array 1000, on thewearable device are transmitted to a computer that communicates witheach detector module 102, as in the case of the other detector arraygeometries already discussed. Wearable detectors may be of particularvalue in situations where involuntary or required patient motiondegrades image quality. For example, a patient could wear a detectorvest 2000(a) or an open frame detector vest 2000(b) while undergoing atreadmill cardiac stress test. Instead of the patient trying to maintainthe same position with respect to the detector vest, 2000(a) or 2000(b),the detector vest 2000(a) or 2000(b) remains aligned with the patient. Apatient can also wear the detector vest 2000(a) or 2000(b) while lyingdown or riding a stationary bike. The ability to utilize detector arraysin situations where involuntary or required patient motion may occuralso applies to wearable detectors other than detector vests (2000(a) or2000(b)), including helmets 2000(c), neck braces or neck wraps (notshown), brassieres 2000(d), and belts/girdles (not shown).

A wearable detector may allow multiple studies, using multipleradionuclide tracers and appropriate detector energy windows, to beconducted at one time. For example, thallium, technetium 99 m, and apositron emitter used together could permit metabolism, regional bloodflow, perfusion abnormalities, and ventricular function to be studiedwhile exercising.

Configurable Detector Arrays

Turning to FIG. 3 a–3 d, movable detector arrays are illustrated. When adetection system involves movable detector arrays 1000(a), the motion ofthe detector arrays 1000(a), or the detector modules 102, do not have tobe confined to a track. Another implementation of the detection systempermits the detector array 1000(a) to move as a unit, but the modules102 within the array can still be positioned independently. This allowsthe array 1000(a) of detector modules 102 to be refocused or individualdetector modules 102, as well as the radiation detectors 101incorporated into those modules 102, to be repositioned as required inorder to optimize detection for a specific type and distribution ofradionuclide. For example, a detector array 1000(a) of fixed or limitedconfigurability can be used in a plane, “L” (FIG. 3 b), or open box(FIG. 3 a) geometry. FIG. 3 a shows perspective views of “box”-shapedimplementations of detector arrays 1000(a) for imaging the heart. Simpleversions include standard geometries that are somewhat fixed. In thisexample, variants of a rectangular shape are possible. Detector arrays1000(a) are present on two/“L”-shaped (FIG. 3 b), three (not shown), orfour (FIG. 3 a) sides of the subject. Each detector array 1000(a) ispositioned by an electronically controlled actuator arm 130.

Turning to FIG. 3 c, a compliant detector array 1000(b) is shown. Thecompliant detector array 1000(b) is comprised of detector modules 102that are small in size, offering more orientation and positioningoptions. The compliant detector array 1000(b) enables the array 1000(b)to follow the approximate contour of the region of interest, as shown inFIG. 3 c, thereby allowing for a contoured geometry for the array1000(b). Positioning may also be accomplished with or without the aid ofactuator arms 30. As illustrated in FIG. 3 c, the flat side of thesubject may be positioned immediately next to a conventional, planardetector array 1000 so that the system does not require an actuator arm.

A gantry that would normally be used for encircling the subject is notnecessary in this embodiment since the flat side of the subject can bepositioned next to the planar detector array 1000. The computer systemmonitors and records in real time the position and orientation of thedetector arrays 1000 and 1000(b), including detectors modules 102,detectors 101, and collimators 105, and relevant motions with respect tothe subject.

FIG. 3 d illustrates a configurable detector system 3000 that utilizesconfigurable arrays 1000(a) manipulated by computer-controlled actuatorarms 130. The actuator arms 130 are able to move along slides or rails131 within a gantry 135, thereby allowing the detector system 3000 to bereconfigured “on the fly”. The actuator arms 130(a) are more flexiblethan the actuator arms 130 shown in FIG. 3 a and 3 b due to theinclusion of rotation capability. The lower detector module array 133 isshown as stationary since the subject is positioned next to it in thisconfiguration. Actuator control of the lower detector module array 133may be implemented if desired, thereby allowing dynamic position of thelower detector array 133 in the same manner as the other detector arrays1000(a) illustrated. Position and orientation information of any of thedetector arrays 1000(a), 133 with respect to the subject is recorded.

FIG. 3 e illustrates one example of configurable arrays 1000(a) mountedon actuator arms 130(a) that are capable of rotational motion andadditionally capable of tilting the arrays 1000(a). The ability to tiltthe arrays 1000(a) with respect to the actuator arms 130(a) makes thearrays 1000(a) adaptable to a greater range of patient geometries. Theoverlap between detector arrays 1000(a) is reduced providing each array1000(a) with its own actuator arm 130(a), thereby enabling the movementof each array 1000(a) independently of the other arrays 1000(a). Theactuator arms 130(a) preferably are capable of bidirectional movementalong a gantry 135. Communications links 103 enables a computer systemto control the actuator arms and the positioning of the arrays 1000(a).

Alternatively, as illustrated in FIG. 3 f, three detector arrays1000(1)–(3) may be joined in operable connection with each other withonly one of the arrays 1000(2) attached to the actuator arm 130(b). Thearrays 1000(1)–(3) are connected such that a middle array 1000(2) isconnected on each end of the array 1000(2) to one end of the otherarrays 1000(1), 1000(3). The arrays 1000(1)–(3) are connected by hinges137 that enable arrays 1000(1) and 1000(3) to be tilted relative toarray 1000(2). The hinges 137 preferably are configured such that thearrays 1000(1) and 1000(3) are able to move in a bilateral directionaway from and towards the middle array 1000(2) while also being able totilt relative to the middle array 1000(2). The middle array 1000(2) isadditionally in operable connection with the distal end of the actuatorarm 130(b). As with the embodiment illustrated in FIG. 3 e, the actuatorarm 130(b) is preferably capable of bidirectional movement along agantry 135. A communications link 103 is provided that enables acomputer system to control the movement of the actuator arm 130(b) andthe arrays 1000(1), 1000(2), and 1000(3). Additionally, arrays 1000(1)and 1000(3) are able to be positioned manually rather than solely bycomputer remote control.

Unconventional Collimators

Turning to FIG. 4 a, a perspective, frontal view of an edge-on detectormodule 102 with an unconventional minifying or tapered capillary x-rayoptic collimator 140 is shown. This collimator 140 has increased theapparent aperture of the detector module 102, thereby enabling thedetector module 102 to detect a larger proportion of the sourceradiation. Additionally, the detection of scattered or non-directionalphotons are decreased. The collimator 140 output face, or aperture, isclosely matched to the actual detector aperture. This configurationallows for the design of the module 102 to remain compact. FIG. 4 bshows a side view of the same detector module 102 with a minifyingcapillary x-ray optic collimator 140. Additionally, an optionalconfigurable x-ray optic collimator 141, which is a configurablemultilayer x-ray mirror that functions as a narrow bandwidth anddirectional filter, is introduced.

Conventional, including coded apertures, collimators and unconventionalcollimators can be used with the present invention. Typical conventionalattenuating, rigid geometry collimating devices range from simplepinholes and slats or septa to focused and unfocused grids.Unconventional collimating devices include x-ray optics, configurablecollimators, and Compton scatter module collimators. Cost and imagingrequirements, including types of detectors, spatial resolution,radiation energies, size of subject, and source distribution andintensity, will influence the selection of collimators. Examples ofx-ray optics include devices such as x-ray mirrors, Bragg crystals,pyrolitic graphite crystals such as those described in Nelson et al.,U.S. Pat. No. 4,958,363, filed Aug. 12, 1988, and Nelson et al., U.S.Pat. No. 4,969,175, filed May 10, 1989, which are hereby incorporated byreference for all they disclose and describe, capillary x-ray optics,refractive x-ray lenses, diffractive x-ray lenses and structures, andx-ray Fresnel lenses. If the x-ray optic device or collimator has theability to focus radiation then it may be used to expand the apparentaperture of the detector, as shown in the module 102 illustrated in FIG.4 a, thereby making more efficient use of the radiation source.Minifying capillary x-ray lenses, refractive x-ray lenses, diffractivex-ray lenses and structures, curved x-ray mirrors, and x-ray zone platesare examples of focusing x-ray optic devices. For typical nuclearmedicine applications the x-ray optic collimator may be integrated withthe detector module 102, as in FIG. 4 a, or held in a separate frame andthen aligned with a detector module 102, as in FIG. 5 b(ii).

An alternative design is to use the dual lens or nested lensconfigurations of FIG. 5 a(i), FIG. 5 a(ii), FIG. 5 b(i), and FIG. 5b(ii) to produce multiple focused beams, but each beam is directed to acorresponding detector module. In this instance multiple slit-like beamsare shown. A similar design has been described previously, but narrowbandwidth x-ray mirrors were described rather than refractive lenses.Nelson et al., U.S. Pat. No. 4,958,363, describes such a design. Thepracticality of utilizing specific x-ray optic devices is influenced byseveral factors including: the size and distribution of the radiationsource, the energy spectrum of the radiation, physical size limitationsof the detectors used in the detector array, the imaging format, whichmay be an internal radionuclide source or an external radiation source,and space requirements, and the cost and maintenance of the x-ray opticdevices. If the narrow bandwidth filtering directional discriminationproperties of a x-ray mirror are not needed or the operational energyrange is better suited to refractive or capillary x-ray optics thenthese devices may be preferred.

Turning to FIG. 5 a(i) and FIG. 5 a(ii), perspective views ofconfigurable dual x-ray refractive lenses 150 that incorporaterefractive slats 153 are shown. This design is similar in principle to adual mirror x-ray telescope. The lens includes a support 152 for therefractive slats 153. This lens increases the apparent aperture of aslit-like or slot-like detector. The focal length of the lens needs tobe accounted for when determining where to locate the lens relative tothe source and detector.

As illustrated in FIG. 5 b(i) and FIG. 5 b(ii), this configuration maybe modified and extended by nesting pairs of x-ray refractive lenses150. The nested pairs of configurable or fixed refractive lenses 150 maybe mounted in an assembly 151 similar to the nested x-ray mirrortelescopes that have been used in x-ray astronomy for a number of years.Practical nested lens will require refractive lenses with thin shellsused as supporting structures. The efficiency of the nesting techniquewill improve if the refractive lenses 150 can be densely packed.Therefore, a relatively thin support structure 151 is preferred.

FIG. 6 shows a perspective view of an electronically-controlledconfigurable collimator 160 which uses electronically-controlledconfigurable elements, in this case two sets of configurable slats 161or septa made from appropriate photon attenuating materials such as tin,lead, tungsten, or uranium. The two sets of adjustable slats 161 includea longer length slat set 163 and a set of shorter length slats 164. Thecollimator 160 includes hinges 162, 167 and a support frame 165 thatfacilitates the manipulation of the slats 161. The configurablecollimator 160 can be implemented so that the slats 161 moveindependently or in synchrony, similar to the operation of windowblinds. The long slats 163 running the length of the detector mayreplaced with slats that are subdivided such that each subdivided slatis individually manipulated for each corresponding detector element. Theshorter slats 164 are configured so that a set of two shorter slats 164defines the edges of a detector module 102, with one short slat 164defining a distal edge and the other short slat 165 defining a proximaledge of the detector module 102. The slats 163, 164 may be manipulatedthrough the use of devices such as actuators, miniature motors, andpulley mechanisms. Moreover, the collimator 160 is capable of beingcontrolled remotely through its communications link 103. Collimator 160may be utilized with the detector modules 102 of any of the detectorarrays 1000 disclosed herein.

Configurable collimators may utilize actuators, includingelectromechanical biopolymers or piezo-drivers, small motors,micromachines, or screw drives, to control parameters such as aperturesize and orientation. In alternative embodiments of the configurablecollimator, refractive, diffractive, or reflective elements aresubstituted for the attenuating elements, thereby forming configurablex-ray optics. An immediate extension of this approach is to abutconfigurable collimators to create 2-D arrays of attenuating,reflective, diffractive, or refractive slats. For example, if reflectiveslats are made sufficiently small these elements will assume the role ofmicromirrors, resulting in highly-configurable micromirror array x-rayoptics which can be electronically-controlled. Optical micromirrors arecommercially available. Additional functionality can be added if eachmicromirror is constructed on an actuator or a deformable surfacecontrolled by actuators. By manipulating the elevation as well as thetilt of the micromirrors, adaptive 3-D x-ray optics can be formed.

A variation of this device is to replace the x-ray mirror coatings withdiffractive or refractive structures. In a similar manner, neutronmirror coatings or refractive or diffractive structures can be used withmicromirrors for neutron radiography applications.

Configurable x-ray optics are useful not only for modifying theradiation incident on the radiation detector but also for modifying theradiation emitted by a radiation source, including a radionuclide orradiographic x-ray source. For example, the configurable dual x-rayrefractive lenses 150 shown in FIG. 5 a(i) and FIG. 5 a(ii) useelectro-active, electromechanical biopolymer actuators to adjust therefractive elements in the lenses.

Turning to FIG. 7 a, a perspective view of an electronically controlledconfigurable refractive lens 150 and a configurable, single elementx-ray mirror 141 are illustrated. The apparatus in FIG. 7 a provides anexample of compound x-ray optics. The refractive slats 153 incorporatean electromechanical biopolymer material that functions as anelectronically controlled actuator, thereby facilitating themanipulation of the refractive slats 153 without requiring a separateactuating mechanism. The tilt of each element can beelectronically-controlled by manipulating the electro-active biopolymermaterial directly.

In FIG. 7 b, the configurable refractive lens 150 is replaced with afixed-focal length capillary x-ray lens 70, which is paired with theconfigurable, single element x-ray mirror 141. The x-ray mirror 141provides spectral and directional filtering and is less expensive tomanufacture than a focused x-ray mirror. If additional focusing isdesired a focused x-ray mirror or a refractive x-ray lens could be usedin place of the single element x-ray mirror 41. If additional spectralor directional filtering is not needed, then the flat x-ray mirror 141can be eliminated.

The simple configurable collimator implementations describe hereinpreferably use electronically-controlled mechanical means to tilt rigidelements or flex elements into the desired position according topredetermined settings or based on detector feedback. More complicatedimplementations preferably use actuators or micromachines to adjust thesurfaces of reflective or refractive elements. The technique of usingdetector feedback to control actuators is well-known in the field ofadaptive optics where non-ionizing electromagnetic radiation istypically employed.

Turning now to FIG. 7 e, another collimator encompassed by the presentinvention is illustrated. A wedge-shaped capillary x-ray optic lens 180is aligned with a very dense array of radiation sources 181. The densearray of radiation sources 181 effectively forms a continuous source ofradiation since the sources are packed in extremely close proximity toeach other. This configuration allows the generation of a slit-like beam182 from the radiation source 181. Here, the capillary x-ray lens 180can be simplified for a slit scan application since focusing is onlyneeded along one dimension. The wedge-shaped capillary lens 180 offersfocusing along one direction, such as along the width of the slit. Ifadequate intensity can be obtained without focusing then a non-focusingcapillary lens may be substituted into this configuration to functionsimply as a highly directional collimator. With this embodiment, arotating cylindrical anode tube could also be used with the extendedfocal spot or the extended focal spot could be simulated using a fastscanning electron beam. Alternatively, other x-ray optic devices may besubstituted for, or used in conjunction with, the capillary x-ray lens180. This technique requires the use of x-ray optics of increasedcomplexity further from the center of the focal spot due to theincreasing angle of incident radiation at the x-ray optics, assuming thegoal is to generate approximately parallel x-rays. An alternative toparallel scanning slits is to use slits in a radial geometry, permittingrotational scanning. This scanning format could be used in mammographywith either flat or curved compression plates.

Another technique for reducing the scan time is to increase the numberof scanning slits. The X-ray optics described in FIG. 5 b(ii) could beemployed for this purpose. As seen in FIG. 5 b(ii), all slits and theirdetectors can share a single focal spot.

Alternatively, each slit can be aligned with its own focal spot. Forexample, a slit scanning system for mammography could use at least twox-ray tubes, i.e., two focal spots, each with a focusing x-ray opticcollimator 70 and an aligned detector array 1000 on the other side ofthe subject, as seen in FIG. 9 b.

FIG. 8 illustrates a novel, unconventional collimator 190 that extendsthe principle of a detector array based on modules to a semiconductorCompton scatter detector array based on modules. This new collimator 190preferably is used for nuclear medicine Compton scatter imaging. The newcollimator 190 is also capable of being used with a standard Gammacamera or an array of detector modules to enhance the current Comptonelectronic Gamma camera design. FIG. 8 shows a perspective view of theCompton scatter module collimator 190 which preferably is a planardetector array comprised of strip detector modules 191, which aredouble-sided, crossed strips for 2-D resolution, positioned within aframe. In this implementation, modules 191 which incorporate relativelythin, linear or 2-D semiconductor detectors can be configured into anumber of geometries compatible with the standard Gamma camera or arrayof detector modules. Relatively thin, linear or 2-D semiconductordetectors will be much less costly to manufacture than the thick,moderately-large, 2-D semiconductor array detectors currently beingtested and are therefore preferred.

The strip detector modules 191 shown in FIG. 8 are double-sided. Theback-side strips (not shown) are oriented at 190 degrees to thefront-side strips, thereby providing 2-D information. Compton scatterphoton radiation is formed from this configuration and the radiation isdetected by a Gamma camera. The Gamma camera preferably is locatedbehind the collimator 190 and is not shown in FIG. 8. A connection 192preferably is provided to transmit output from each detector module 191to signal processing electronics. Module parameters, such as, e.g.,temperature, electronic readout, and power are electronically controlledvia the connection 192. In one embodiment, the modules 191 are alsocapable of interfacing with other modules 191 in order to shareresources such as power, cooling, and communications. Fixed orconfigurable collimator geometries of greater complexity can beimplemented as needed. A supporting frame 193 preferably is provided inorder to maintain the positioning of the detector modules 191.

In one embodiment, the detectors modules 191 are immersed in a lowtemperature coolant in order to prevent operating temperatures that aretoo high. In this embodiment, all of the modules 191 can be encapsulatedtogether in a container that holds the coolant. The collimator 190 canalso be dithered to compensate for dead spaces between detectors thatare closely spaced.

Imaging Systems

The present invention may also be employed in x-ray radiographic imagingsystems. Operational energy ranges and spatial resolution requirementswill be different in many instances from those that are used in nuclearmedicine. Two additional factors that impact the design of nuclearmedicine and x-ray radiography imaging systems are the image acquisitiontime and the properties of the radiation source. A consequence of theshort acquisition times required in x-ray radiography, which arefractions of a second to multiple seconds for 3-D imaging, is the needfor an intense radiation source or a subject who is not heavilyattenuating. Some nuclear medicine study scan times can exceed 15minutes. Typically both nuclear and x-ray radiographic imagingmodalities benefit from an increase in the apparent intensity activityof the source. Benefits include reduced acquisition times and/orimproved statistics. The source distribution and location are poorlydefined in many nuclear medicine imaging applications, limiting thevalue of customized focused photon optics for directing more of thesource radiation to a detector. In contrast to this situation, thesource distribution, which is highly localized, and position are usuallywell-defined in x-ray radiography. Focusing x-ray optic collimators canbe designed for a specific x-ray tube focal spot distribution. Thiswould not be practical given cost constraints for most nuclear medicineimaging applications. X-ray radiography applications that could use oneor more detector modules include slit, slot, or CT scanning.

X-ray mammography is a radiographic imaging application that usesrelatively low x-ray energies, increasing the number of viable detectorand unconventional collimator choices. For example, the source shown inFIG. 7 b can be the focal spot of a x-ray tube. The tube is preferably asource with a well-defined location and reasonably well-defineddirectional and spectral properties. The focused capillary x-ray lens70, or a refractive lens, a diffractive lens, a curved or configurablex-ray mirror, nested lens, or combinations of x-ray optic collimators,would be used to increase the intensity of radiation that wouldultimately be detected by a slit-shaped detector, such as an edge-ondetector, after passing through the subject. This configuration isillustrated in FIG. 9 a.

Gantry Imaging Systems

Turning to FIG. 9 a, a perspective view of a rotatable gantry unit 1200with an adjustable arm 1100 configured to hold an x-ray tube 1102,incorporating a radiation source 111 and a capillary x-ray lens 70, anda detector array 1000 is shown. The array 1000 may use either analog ordigital detector modules. The gantry system 1200 rotates about an axisand has an arm 1100 that allows for further adjustment in abidirectional manner. As shown, the gantry system 1200 is adapted forx-ray mammography applications by incorporating a pair of compressionplates 1101 that are used to position a subject breast. This design iscomparable to a traditional x-ray film-screen mammography-imaging unitwhich utilizes a rotatable gantry. The x-ray tube 1102 is aligned withthe detector array 1000. The x-ray tube 1102 and detector array 1000 arethen scanned as a unit.

As seen in FIG. 9 c, if additional spectral and directional filtering ofthe radiation beam is desired then a configurable x-ray mirror 1110 canbe inserted between the capillary x-ray lens 70 and the compressionplates 1101 holding the subject compressed breast 1111, as seen in FIG.9 c. A second x-ray mirror or crystal may be positioned between thecompression plates 1101 and the detector array 1000 if even morespectral and directional filtering is desired. In some instances arefractive x-ray lens can be used in place of the x-ray mirror 1110,providing focusing and some filtering instead of the spectral anddirectional filtering provided by a configurable x-ray mirror.Collimators 1112 are provided on either side of the compression plates1101 in order to limit the x-ray beams that are passed through thesubject breast 1111. The collimators 1112 concentrate a segment of thex-ray source output into a slit or slot geometry.

Turning now to FIG. 9 b, a dual gantry system 1300 that combines twoindividual gantry arms 1110 is illustrated. The separation distancebetween the gantry arms 1110 can be adjusted for scanning objects, forexample, compressed breasts, of various sizes. As with the gantry system1200 shown in FIG. 9 a, the dual gantry system 1300 may also be rotatedand the arms 1110 are adjustable in a bidirectional manner. Compressionplates 1101 that are typically employed in mammography imaging may beincorporated into the dual gantry system 1300 but are not shown in FIG.9 b. An x-ray tube 1102, incorporating a radiation source 111 and acapillary x-ray lens 70, and a detector array 1000 are attached to eacharm 1110 and are adjustable by virtue of being mounted on the arms 1100.Each unit would only be required to scan half as far as a single unitsystem, reducing the total scan time and x-ray tube operational time by50%.

Many gantry designs are possible, including portable gantries designs,which are in use with portable x-ray and Gamma camera imaging systems.These gantry designs allow improved detector positioning with respect tothe heart in comparison to standard Gamma camera designs.

Composite Anodes

The present invention is additionally directed to the generation offocused radiation by operating a composite anode operating inconjunction with x-ray optic lenses. Turning to FIG. 7 c, a compositeanode 73 comprised of N types of disks is illustrated. In this case Nequals 2 and the disks preferably are molybdenum and rhodium, although Nequals 1 or N greater than 2 can be constructed. For example, the disksare capable of being manufactured using other materials, such as, e.g.,tungsten. Alternatively, another material may be added to the primarymaterial used to manufacture the disk.

An array of capillary x-ray optic lenses 71 preferably is aligned withan array of radiation sources/focal spots 72(a) that project multipleelectron beams incident to a rotating cylindrical anode to form anextended slit-like radiation beam generated by an extended radiationsource. This configuration allows the radiation to be focused into theextended slit radiation beam. As discussed, the composite cylindricalanode 73 shown in FIG. 7 c preferably is comprised of rhodium disks 74and molybdenum disks 75. Shifting the tube laterally by one disk widthwhile maintaining the positions of the focal spots 72 and capillaryx-ray lens array 71 permits the selection of a specific anode spectrum.In the configuration shown in FIG. 7 c, the choice is between a spectrumgenerated by the rhodium disks 74 and a spectrum generated by themolybdenum disks 75. Coolant 76 is passed through the anode 73 in orderto facilitate the maintenance of an optimal operating temperature.

Other composite anode tube geometries are possible. For example, ananode could be built by combining 2 or more fractional, including halfor quarter, circle cylinders. FIG. 7 d illustrates an oscillatingcomposite anode 79 comprised of two materials, each in the shape of afractional circle cylinder. A plurality of focal spots 72(b) is locatedlongitudinally on each fractional cylinder 77, 78. Two materials thatare capable of being used to manufacture the fractional cylinders 77, 78include rhodium and molybdenum. The configuration shown in FIG. 7 d, asthe configuration in FIG. 7 c, allows for a choice of spectrumsgenerated by the materials comprising the oscillating composite anode79.

Another alternative geometry is evidenced in a continuous slit scanacquisition format. In a continuous slit scan acquisition format, theanode would oscillate through an arc slightly smaller than the fractionof a circle which the desired material, such as molybdenum, rhodium, ortungsten, occupies. In this case only a single material is used todetermine the x-ray spectrum distribution.

Another alternative embodiment uses at least two full-sized anodecylinders comprised of different materials which can be shifted in andout of position so that the same electron beam source can be used. Onlythe anode that is being used to generate x-rays needs to rotate. Thisparticular multiple-anode configuration is used so that the x-ray tubeunit will remain reasonably compact.

Yet another alternative embodiment is to abut different and distinctanodes such that a single elongated anode is present. With thisembodiment, an entire anode can be shifted depending on which anodematerial is desired.

This approach can be extended to the deployment of multiple focusingcapillary x-ray lenses so that multiple slits or slots can be scanned atthe same time using a single x-ray tube focal spot. A drawback to theuse of a capillary lens to focus a small focal spot onto an extendedslit is that the capillary x-ray optics will become more complex inorder to direct radiation, in an approximately parallel direction, tosections of the slit which are far from the center of the slit. Thecapillary x-ray lens design can be simplified if the focal spot sourceor radiation source is reshaped to more closely match the shape of theslit or slot. The x-ray source can be modified such that additionalfocal spots are incorporated parallel to the length of the slit. Anequal number of capillary x-ray lenses or a lens array with relaxeddesign constraints are abutted and aligned with the corresponding focalspots in the focal spot array of a rotating cylindrical anode x-raytube.

Compression Plates

It should be noted that the use of compression plates to compress thearea of the subject being imaged may be desirable. A reasonably smalltissue path between the radiation source and the detector device ishighly desirable in both nuclear medicine scintimammography and x-raymammography. This reduces absorption, scatter, and in general improvesimage quality. In scintimammography partial, i.e., limited, breastcompression can be used, allowing the detector device to move a fewcentimeters closer to a potential tumor. By comparison, more-strenuouscompression is applied in standard x-ray mammography since the breast isunder tension for a much shorter period of time. The sensitivity andspecificity of scintimammography can be improved by using theinformation acquired in the initial scan to re-image suspicious regions.Re-imaging involves applying increased compression to a smaller sectionof the breast using contoured or flat compression plates of reducedarea. Contoured and flat compression plates of reduced area, includingversions with an open region in the compression plate, for opticalimaging of tissue have been described by Nelson et al., U.S. Pat. No.5,999,836, filed Feb. 2, 1996, which is fully incorporated herein byreference for all it discloses and describes. In a compression platewith an open region, the open region is located adjacent to the skinsurface. This open region typically allows air and/or a couplingfluid/gel to be in contact with the skin surface. Radiation from anacoustic source can also be coupled into and out of the open region(s),enabling compression transmission and backscatter (reflection) acousticimage data and acousto-optic image data to be acquired as well asoptical image data. The size and geometry of the open region in acompression plate can be customized according to the application.

A compression plate of reduced area refers to the actual plate surfacewhich is used to compress a section of the breast to a uniform thicknessrelative to the surface of a typical compression plate which is used tocompress the entire breast to a uniform thickness. For example, if astandard compression plate is translated with respect to the center ofthe breast such that approximately 50% of the breast are compresseduniformly, then this functions as a compression plate of reduced area.This additional level of compression permits a detector module array tobe positioned nearer to a potential tumor. If a hole is included in thecompression plate even closer positioning is possible while allowingother instruments, including instruments for ultrasound imaging, foroptical imaging, and for injecting materials or obtaining tissuesamples, to gain access to the compressed area. In some instances it ispossible to forego the initial imaging of the entire breast and insteadbegin by imaging smaller sections, which are compressed with increasedforce relative to compression of the entire breast, if the total imageacquisition time is acceptable. In this image acquisition formatadjacent images should have sufficient overlap so that potentialstructures of interest will be visible in at least one image section.

The concept of compressing a smaller section of the breast morestrenuously than would be tolerated for whole-breast compression inscintimammography in order to achieve greater local compression can beapplied to x-ray mammography. One rather limited approach is to positionthe patient such that the left edge or right edge of a standardcompression plate, approximately 24–30 cm×18–24 cm, is near the centerline on the breast, slightly more than one-half of the breast could becompressed and scanned. A x-ray technologist would then reposition acompression plate relative to the breast such that the other half iscompressed and scanned. If higher levels of compression are desired thenone or both flat compression plates need to be reduced in size.

Contoured or flat compression plates, including plates with holes, thatcan compress only a section of a breast rather than the entire breast,as is practiced in x-ray film-screen mammography, may eliminate the needor simplify the requirements for items such as x-ray optics, multiple orextended focal spot sources, and multiple slits. X-ray tube powerhandling requirements could be reduced since continuous scanning occursacross a smaller area in comparison to conventional x-ray mammographyimaging even if the compressed tissue thickness were to remain the same.Additionally, greater levels of compression can be attained if one orboth of the plates of reduced size are contoured.

Turning to FIG. 10 a, a perspective view of a contoured uppercompression plate 1120 that is appropriate for compressing a section ofa breast 1111 is shown with the image scan area 1123 indicated. In thisembodiment, the bottom compression plate 1121 is flat, simplifying thepositioning of the breast 1111.

In FIG. 10 b, the flat bottom compression plate is replaced by a secondcontoured compression plate 1122. This configuration enables additionalcompression of the breast 1111 as compared with the configurationillustrated in FIG. 10 a.

The present invention is also directed to a method of acquiring a seriesof overlapping successive sections or sub-images in order to increaseimage resolution of the subject. Turning to FIG. 10 c, the overlap 1133that is produced by acquiring successive image sections 1130, 1131, 1132of the subject are illustrated. This overlap 1133 can be utilized toproduce a continuous, higher resolution image. The overlap 1133 must besufficient so that small structures 1134, 1135 may be viewed in at leasta single image area. To acquire the successive image sections 1130,1131, 1132, after each section is scanned, the compression plate orplates are repositioned relative to the breast before the next sectionis scanned. Efficient repositioning involves marking temporary spots onthe breast or using a focused light beam on the breast in order todefine the locations of the sections to be compressed. The compressionplate or plates are repositioned such that a small overlap betweenadjacent sections occurs. This is continued until the entire breast isscanned. A series of high resolution breast section images can then beevaluated by a radiologist.

Section areas should be sufficiently large such that any structures ofinterest can be clearly discerned within at least one section image.Advantages include lower patient dose and higher spatial and contrastresolution, less stress on the x-ray source, reduced scatter, and theoption to use a less-energetic x-ray spectrum. Additionally, if theprocedure is video recorded, the image sections can be referenced to avideo recording of the various positions of the compression plates inorder to increase the accuracy of the scans. An alternativeimplementation of this technique is to acquire a complete scan initiallyand then selectively re-image problematic sections using increasedcompression.

Dynamic Acquisition of Images

The present invention is also directed to a process for dynamicallyacquiring partial images of a subject image in order to obtain anentire, optimized image of the subject. This method is particularlyadvantageous when utilized to acquire images of a breast. The process ofimage optimization based on the energy spectrum and integrated intensitywhile limiting patient risk is complicated by the fact that breasts aretypically non-uniform in tissue composition and the tissue distributionsare non-uniform. In order to determine a reasonable compromise x-rayspectrum it would be desirable to implement a static or dynamic pre-scansince both techniques permit dynamic acquisition of the mammographyimage. Image optimization is obtained by scanning a tissue volume thatis no finer than the area of the slit or slot. A static pre-scanacquires an entire image at a low radiation level in order to determinethe degree of attenuation while avoiding the radiation levels needed toacquire an acceptable image. After the pre-scan, the actual image isacquired based on all of the pre-scan data. In this case the tubevoltage and current could be dynamically controlled during imageacquisition. If the tube uses an array of radiation sources (see FIG. 7e) then individual sources or sub-arrays of sources can be configured asneeded in order to provide the desired beam characteristics during ascan. A dynamic pre-scan uses two slits or slots which move in parallel.The first slit or slot is used to acquire the low radiation level datawhich is then used to dynamically and adaptively adjust the x-ray tubecurrent and/or KV so that an acceptable signal-to-noise ration (SNR) ismaintained for each tissue segment imaged by the second slit or slot.

Alternatively, the pre-scan may be avoided by adjusting the intensity,i.e., the current, and/or tube voltage dynamically (“on-the-fly”) inorder to maintain an acceptable SNR. A feedback system manipulates beamcurrent and/or KV. At each slit or slot position, the current level isinitially reduced and the detected output is analyzed. The beam currentand/or tube KV is then increased to the appropriate level or the timethe slit takes to scan a slit/slot area is increased so as to acquireadequate statistics. Alternatively, both the beam current and the scantime may be increased in order to acquire statistics. Another,less-complex, dynamic acquisition technique is to operate with aconstant beam current and then track the time needed to acquire adequatestatistics for each slit or slot area.

Single-slit and multi-slit designs are known to those skilled in theart-of x-ray radiography. The utility of such designs can be enhanced byincorporating x-ray optics with traditional or novel focal spotconfigurations in conjunction with efficient detector modules.Additional benefits are gained in x-ray mammography andscintimammography by modifying the scanning procedure so that increasedcompression is implemented. Preferably, compression plates such as thosedescribed above and in FIG. 10 a and FIG. 10 b are utilized to implementthe increased compression.

Correction and Tuning of an Array

The present invention is also directed to a novel method of correctingand tuning a detector array that preferably involves tracking one or alimited number of spheres, such as microspheres, or small capsules, saidspheres or capsules containing known levels of radioactivity, as theyare taken up by or circulate within the patient. These are collectivelyreferred to as reference sources. Since the sizes, compositions,activities, and photon energies of the reference sources are known orare measured prior to their introduction into the patient, thescattering and absorption effects of tissue positioned between thesereference sources and the detector can be measured directly once thereference sources are at the desired locations. The reference sourcescan be designed to be biodegradable or inert depending on how long orwhere they are expected to be within the body. The reference sources canhave internal structures and non-uniform activity distributions.Typically the reference sources are introduced into the patient prior tothe nuclear medicine test. The reference sources can also function asdistinct, internal, small sources that can be measured with little or nointerference from other reference sources. Individual reference sourcescan have distinctive properties such as different levels ofradioactivity, different types of radionuclides, or the incorporation ofmagnetic, acoustic, inductive, or x-ray attenuating materials. Theseproperties can be useful for identifying specific reference sourceswithin the body, measuring the effects of tissue at different energies,helping to guide the reference source to a desired location, andproviding position information by causing the reference source toabsorb, reflect, or emit acoustic, EM, or ionizing radiation wheninterrogated by an external field. The reference sources can also betracked as they move within the body. Once the relatively smallreference sources are in the appropriate locations they also can be usedto fine tune and focus the detector array for that specific imagingtask. Thus, the position-dependent imaging capabilities of the detectorarray can be estimated for a patient. If the reference sources aresufficiently distinct from the radionuclides introduced into the patientduring a nuclear medicine test then estimating attenuation and focusingcan be done dynamically, permitting adaptive imaging.

The concept of adaptive imaging is utilized in many imagingapplications. In particular, the use of artificial guide-stars iswell-known in Astronomy. An important difference between our referencesource and a guide-star is that the intensity of the guide-star is notparticularly important. The guide-star is used to correct phasedistortions to an optical wavefront due to a turbulent atmosphere. Themethod of the present invention attempts to estimate attenuationcorrections and use the reference source to help focus the detectorarray at approximately the position of the actual radionuclidedistribution. The correction and tuning/focusing method along withappropriate reference sources can also used with existing Gamma cameras,PET scanners, etc.

The present invention is additionally directed to a process ofcalibrating a detector module array using a known source distribution.Electronic calibration of a detector module array involves using a knownsource distribution such that the responses of individual detectormodules can be balanced either electronically or through softwareamplification of the digitized data. This calibration effort willinclude evaluating detection events that are recorded by more than onedetector module, which is similar to the Gamma camera problem ofevaluating detection events recorded by multiple PMTs.

Typical source distributions include collimated spots, slits, slots, orflat fields with appropriate energy distributions. It is assumed thatsource energy distribution is appropriate for the imaging task thedetector modules will be used for. If the detector offers energyresolution then an additional calibration can be performed to accountfor energy resolution. The energy-dependent Modulation Transfer Function(MTF (E)) can be measured over the expected energy range of the x-raysource or from a series of measurements involving narrow band sourceswith different energies.

The process may be applied to the radiographic imaging application ofx-ray mammography. The x-ray source properties are well-defined. Intraditional x-ray film-screen mammography using an integrating detector,the intensity of the x-ray field decreases or falls off as the positionchanges from the center to the edge of the x-ray field. The result is adivergent beam from the focal spot, said focal spot being a point-likex-ray source. Calibration of this relatively large, planar, x-ray fieldis typically not done. Next consider replacing the film-screen detectorwith a detector array comprised of a single detector module which isappropriate for slit-scan imaging. Once the detector module is alignedwith respect to the x-ray source then a calibration can be performedthat approximately corrects for the variations of the x-ray beamintensity at the locations of the detector module detector pixels. Thisresults in a position-dependent, energy-dependent, intensity profile.For example, if the slit-like detector module uses an edge-on detector,the intensity of the slit-like x-ray beam along the length of thedetector can be measured. The spectral distribution of a typicalMo-anode or W-anode mammography x-ray tube is relatively broad band,usually greater than 10 KeV. This implies that the information contentof detected photons at the upper and lower extremes of the spectral bandcan be substantially different for the typical x-ray energies used infilm-screen mammography. If the edge-on detector is capable of providingsufficient energy resolution, such as when an energy-resolving detectorrather than an integrating detector is utilized, then additionalinformation is available. Each detected photon represents theexponential attenuation properties of the filter, which in the case ofmammography is breast tissue. The filter, due to its attenuationproperties, modifies the local x-ray beam intensity and spectraldistribution at each detector pixel. If the spectral distribution isuniform along the length of the detector then a reasonable comparison ofcorrected intensity and spectral content between individual pixels inthe detected image can be made. What is essentially acquired is a set ofoverlapping energy-dependent images. If energy-dependent MTF (MTF (E))measurements are available, then an improved analysis of theenergy-dependent images is possible. If the spectral distribution of thesource is not uniform along the length of the detector then theposition-dependent, source spectral intensity distribution can bemeasured and used to approximately correct the detected data.

An alternative embodiment of this process involves narrowing the x-raybeam bandwidth about an appropriate energy for a particular breast-type,adjusting for size and composition. This modification simplifies thedetection and analysis process for both integrating detectors andenergy-resolving detectors. Configurations of detectors that would beappropriate for use in this embodiment of the process are describedabove and in FIG. 5 a(i)–(ii), FIG. 5 b(i)–(ii) and FIG. 7 a–7 d.

It is desirable to measure the detector MTF (E) at the appropriateenergy. In addition, a narrow bandwidth filter may be utilized in orderto reduce patient risk by removing radiation energies which are totallyabsorbed by the breast or represent relatively little information aboutthe properties of breast tissue.

Although the embodiments of the present invention have been described interms of its use for nuclear medicine and x-ray mammographyapplications, the present invention may also be used for other medicalradiographic imaging applications as well as industrial and scientificapplications. For example, similar designs can be used with appropriateradiation collimators such as neutron mirrors or electron optics forimaging sources of neutrons or charged particles, respectively.Radiological and non-medical applications which utilize compositionanalysis based on Compton scatter measurements and/or tomographicimaging will also benefit from this design. Unconventional collimatorssuch as x-ray optic, configurable, and Compton scatter collimators canbe used with standard Gamma cameras to improve the capabilities of thesedevices.

While the invention is susceptible to various modifications andalternative forms, specific examples thereof have been shown by way ofexample in the drawings and are herein described in detail. It should beunderstood, however, that the invention is not to be limited to theparticular forms or methods disclosed, but to the contrary, theinvention is to cover all modifications, equivalents, and alternativesfalling within the spirit and scope of the appended claims.

1. A method of calibrating a radiation detection system comprising:providing a radiation source that emits radiation, wherein the source ischosen from the group consisting of a uniform point-like source, aline-like source, a spherical source, a rod-like source, a collimatedspot source, a slit source, a slot source, a grid pattern source, aplanar flood field, and a shaped three-dimensional flood field,measuring an energy-dependent modulation transfer function of thedetection system, and measuring the level of radiation emitted from thesource that is detected by the detection system, and calibrating thedetection system by evaluating the detected radiation and balancing thesystem based upon the detected radiation and the energy-dependentmodulation transfer function of the detection system.
 2. A method ofestimating the effects of tissue attenuation on the intensity and energydistribution of an x-ray beam comprising: calibrating anenergy-resolving detector array by determining its energy-dependentmodulator transfer function, aligning the calibrated energy-resolvingdetector array with the x-ray beam, measuring a firstposition-dependent, energy-dependent intensity profile of the x-ray beamat the detector array, transmitting the x-ray beam through a patient,measuring a second position-dependent, energy-dependent intensityprofile of the x-ray beam at the detector array immediately after thebeam has been transmitted through the patient, and comparing the firstand the second position-dependent, energy-dependent intensity profilesof the beam.